Conventionally, radiation images such as X-ray images are widely used to diagnose symptoms in medical practice. Particularly, radiation images using an intensifying screen-film system are still used in medical practices worldwide, as an imaging system exhibiting both high reliability and excellent cost performance, as a result of achievement of high sensitivity and high image quality in a long history. However, such image information is so-called analog image information that cannot be subjected to flexible image processing and cannot be instantaneously electrically transmitted, unlike digital image information currently continuing to be developed.
Recent years have seen the appearance of digital radiation detectors typified by radiation detectors of computed radiography (CR) type, flat panel type (flat panel detectors: FPD),and the like. In these radiation detectors, a digital radiation image can be directly acquired and can be directly displayed on an image display device such as a cathode-ray tube panel or a liquid crystal panel, thereby making it unnecessary to form an image on a photographic film. Consequently, such digital radiation detectors, for example, X-ray detectors, reduce the necessity of image formation by a silver salt photographic system, thus greatly improving convenience in diagnostic work in hospitals and clinics.
One of digital technologies relating to X-ray images is computed radiography (CR), which is currently used in medical facilities. However, X-ray images acquired by CR are insufficient in sharpness and spatial resolution as compared to those acquired by screen-film systems such as a silver salt photographic system, and the image quality level thereof is still far from that of screen-film systems. Thus, as another new digital X-ray imaging technology, for example, flat X-ray detectors (Flat Panel Detectors: FPDs) using thin film transistors (TFTs) have been developed.
FPDs as mentioned above convert an X-ray to visible light by using, in principle, a scintillator panel including a phosphor layer made of an X-ray phosphor having properties that convert an applied X-ray to visible light to emit light. Radiography using a low-dose X-ray source requires use of a scintillator panel high in light emission efficiency, which is conversion efficiency from an X-ray to visible light, in order to improve a ratio of signal to noise (a S/N ratio) detected from the scintillator panel. In general, the light emission efficiency of a scintillator panel is determined by the thickness of a phosphor layer and the X-ray absorption coefficient of a phosphor. As the thickness of the phosphor layer increases, emission light generated in the phosphor layer by X-ray irradiation scatters more easily in the scintillator layer, thereby reducing sharpness of an X-ray image that is acquired via the scintillator panel. Accordingly, setting a level of sharpness necessary for image quality of an X-ray image naturally leads to the determination of a limit to the film thickness of the phosphor layer in the scintillator panel.
As used herein, the term phosphor is also referred to as scintillator, and the term “phosphor layer” is also referred to as scintillator layer.
Additionally, the shape of the phosphor forming the phosphor layer is also important in obtaining a scintillator panel that can provide X-ray images having high brightness and high sharpness. Many scintillator panels employ a phosphor having a columnar crystal shape as a phosphor forming a scintillator layer, and are usually formed by arranging a plurality of such columnar crystals on a substrate, a support body, or the like. As used herein, each of the columnar crystals forming the scintillator layer has a shape extending perpendicularly to a main surface of the substrate, the support body, or the like so that each columnar crystal can efficiently discharge emission light generated in the scintillator layer in the direction perpendicular to the main surface thereof. Scintillator panels employing such a layout in the scintillator layer can maintain the brightness and sharpness at high levels, as well as high strength in the direction perpendicular to the substrate, the support body, or the like. The phrase “the direction perpendicular to the substrate, the support body, or the like” may be hereinafter referred to as “film thickness direction”.
In recent years, various studies and attempts have been made to focus on a crystal shape of a phosphor forming a scintillator layer.
For example, Patent Literature 1: WO 2010/032503 discloses a radiation conversion panel in which a phosphor layer including a phosphor matrix that forms columnar crystals having a specific shape are provided on a substrate, as a radiation conversion panel that can obtain X-ray images having high brightness and high sharpness. In the radiation conversion panel described in Patent Literature 1, the phosphor layer includes a first phosphor layer made of the phosphor matrix and having a film thickness within a specific range and a second phosphor layer including the phosphor matrix and an activator. Then, Patent Literature 1 teaches that high sharpness is obtained when, in the columnar crystals of the phosphor forming the phosphor layer, a ratio of a crystal diameter at a height of 10 μm from a substrate side to a crystal diameter at an outermost surface of the scintillator layer is within a specific range.
However, even the radiation conversion panel described in Patent Literature 1 still has room for improvement in terms of the sharpness of radiation images to be obtained.
Specifically, the radiation conversion panel described in Patent Literature 1 defines a ratio of an average circle-equivalent diameter “a” at the position of 10 μm from the substrate to an average circle-equivalent diameter “b” at the outermost surface in the columnar crystals of the phosphor layer, but defines nothing about a range of below 10 μm. Thus, sharpness of radiation images to be obtained is still likely to be further improved by allowing the shape, properties, or the like of the columnar crystals to be more precisely controlled in the region of below 10 μm in the columnar crystals.